X-ray tomographic imaging, in its simplest expression, is an X-ray beam traversing an object, and a detector relating the overall attenuation per ray. The attenuation is derived from a comparison of the same ray with and without the presence of the object. From this conceptual definition, several steps are required to properly construct an image. For instance, the finite size of the X-ray generator, the nature and shape of the filter blocking the very low energy X-rays from the generator, the details of the geometry and characteristics of the detector, and the capacity of the acquisition system are all elements that affect how reconstruction is performed.
In one of many possible geometries, an X-ray source on top of the graph shown in FIG. 1 is emitting an X-ray beam forming a fan, traversing the object. While a wide range of values can exist, typically, the distance “C” is around 100 cm, “B” is around 60 cm, and “A” is around 40 cm. The principle of tomography requires that each point of the object is traversed by a collection of rays covering at least 180 degrees. Thus, the entire X-ray generator and detector assembly will rotate around the patient. Mathematical considerations show that the tomographic conditions are met when a scan of 180 degrees plus the fan angle is performed.
Conventional X-ray detectors integrate the total electrical current produced in a radiation sensor, and disregard the amplitude information from individual photon detection events. Since the charge amplitude from each event is proportional to the photon's detected energy, this acquisition provides no information about the energy of individual photons, and is thus unable to capture the energy dependence of the attenuation coefficient in the object.
On the other hand, semiconductor X-ray detectors that are capable of single photon counting and individual pulse-height analysis may be used. These X-ray detectors are made possible by the availability of fast semiconductor radiation sensor materials with room temperature operation and good energy resolution, combined with application-specific integrated circuits (ASICs) suitable for multi-pixel parallel readout and fast counting.
One major advantage of such photon-counting detectors is that, when combined with pulse-height analysis readout, spectral information can be obtained about the attenuation coefficient in the object. A conventional CT measures the attenuation at one average energy only, while in reality, the attenuation coefficient strongly depends on the photon energy. In contrast, with pulse-height analysis, a system is able to categorize the incident X-ray photons into several energy bins based on their detected energy. This spectral information can effectively improve material discrimination and target contrast, all of which can be traded for a dose reduction to a patient.
Using such photon-counting detectors for medical CT applications requires very high X-ray flux in most CT tasks. In a routine CT scan, as many as 108 photons, or even more, can hit one detector element every second. Accordingly, photon-counting detectors suffer from count loss under high rate X-ray irradiation, e.g., in a clinical CT scan. Photon count loss may occur due to, e.g., detector crystal polarization or pulse pileup.